1. Field of the Invention
The present invention relates generally to the field of biomedical analysis and treatment, and in one exemplary aspect to an apparatus and method for evaluating the need for, and applying, defibrillating electrical current.
2. Description of Related Technology
Noninvasive estimates of cardiac output (CO) can be obtained using impedance cardiography. Strictly speaking, impedance cardiography (ICG), also known as thoracic bioimpedance or impedance plethysmography, is used to measure the stroke volume of the heart. As shown in Eqn. (1), when the stroke volume is multiplied by heart rate, cardiac output is obtained.CO=stroke volume×heart rate.  (1)The heart rate is obtained from an electrocardiogram. The basic method of correlating thoracic, or chest cavity, impedance, ZT(t), with stroke volume was developed by Kubicek, et al. at the University of Minnesota for use by NASA. See, e.g., U.S. Reissue Patent No. 30,101 entitled “Impedance plethysmograph” issued Sep. 25, 1979, which is incorporated herein by reference in its entirety. The method generally comprises modeling the thoracic impedance ZT(t) as a constant impedance, Zo, and time-varying impedance, ΔZ (t), as illustrated schematically in FIG. 1. The time-varying impedance is measured by way of an impedance waveform derived from electrodes placed on various locations of the subject's thorax; changes in the impedance over time can then be related to the change in fluidic volume (i.e., stroke volume), and ultimately cardiac output via Eqn. (1) above.
Despite their general utility, prior art impedance cardiography techniques such as those developed by Kubicek, et al. have suffered from certain disabilities. First, the distance (and orientation) between the terminals of the electrodes of the cardiography device which are placed on the skin of the subject is highly variable; this variability introduces error into the impedance measurements. Specifically, under the prior art approaches, individual electrodes 200 such as that shown in FIGS. 2a and 2b, which typically include a button “snap” type connector 202, compliant substrate 204, and gel electrolyte 206 are affixed to the skin of the subject at locations determined by the clinician. Since there is no direct physical coupling between the individual electrodes, their placement is somewhat arbitrary, both with respect to the subject and with respect to each other. Hence, two measurements of the same subject by the same clinician may produce different results, dependent at least in part on the clinician's choice of placement location for the electrodes. It has further been shown that with respect to impedance cardiography measurements, certain values of electrode spacing yield better results than other values.
Additionally, as the subject moves, contorts, and/or respirates during the measurement, the relative orientation and position of the individual electrodes may vary significantly. Electrodes utilizing a weak adhesive may also be displaced laterally to a different location on the skin through subject movement, tension on the electrical leads connected to the electrodes, or even incidental contact. This so-called “motion artifact” can also reflect itself as reduced accuracy of the cardiac output measurements obtained using the impedance cardiography device.
A second disability associated with prior art impedance cardiography techniques relates to the detection of a degraded electrical connection or loss of electrical continuity between the terminals of the electrode and the electrical leads used to connect thereto. Specifically, as the subject moves or sweats during the measurement, the electrolyte of the electrode may lose contact with the skin, and/or the electrical leads may become partially or completely disconnected from the terminals of the electrode. These conditions result at best in a degraded signal, and at worst in a measurement which is not representative of the actual physiological condition of the subject.
Another significant consideration in the use of electrodes as part of impedance cardiographic measurements is the downward or normal pressure applied to the subject in applying the electrode to the skin, and connecting the electrical leads to the electrode. It is desirable to minimize the amount of pressure needed to securely affix the electrode to the subject's skin (as well as engage the electrical lead to the electrode), especially in subjects whose skin has been compromised by way of surgery or other injury, since significant pressure can result in pain, and reopening of wounds.
It is also noted that it is highly desirable to integrate cardiac output measurement capability into a compact, rugged, and efficient platform which is readily compatible with different hardware and software environments. The prior art approach of having a plurality of different, discrete stand-alone monitors which include, for example, a dedicated, redundant display and/or other output or storage device is not optimal, since there is often a need to conserve space at the subject's bedside or even in their home (e.g., in outpatient situations), as well as cost efficiency concerns. Furthermore, a plurality of discrete stand-alone monitors necessarily consume more electrical power (often each having their own separate power supplies), and require the subject or clinician to remain proficient with a plurality of different user interface protocols for the respective monitors. In many cases, the individual stand-alone monitors are also proprietary, such that there is limited if any interface between them for sharing data. For example, where two such monitors require a common parametric measurement (e.g., ECG waveform or blood pressure), one monitor frequently cannot transmit this data to the other monitor due to the lack of interface, thereby necessitating repeating the measurement.
Recognizing these deficiencies, more recent approaches have involved the use of modular devices, wherein for example a common monitor/display function is utilized for a variety of different functional modules. These modules are generally physically mounted in a rack or other such arrangement, with the common monitor/display unit also being mounted therein. A common power supply is also generally provided, thereby eliminating the redundancy and diversity previously described. However, heretofore, impedance cardiography equipment has not been made in such modular fashion, nor otherwise compatible with other modular devices (such as blood pressure monitoring or ECG equipment), such that such other signals can be obtained directly or indirectly from these devices and utilized within the ICG apparatus. The quality or continuity of these signals, whether obtained directly from the subject being monitored or from other modules, has not been readily and reliably provided for either.
Typical patient monitors include modules for several physiologic measurements such as ECG, blood pressure, temperature, and arterial pulse oximetry. The addition of ICG provides the physician with additional useful clinical information about the patient.
Furthermore, prior art ICG devices (modular or otherwise) do not provide the facility for direct transmission of the data obtained from the subject, or other parameters generated by the ICG device after processing the input data, to a remote location for analysis or storage. Rather, the prior art approaches are localized to the bedside or monitoring location. This is a distinct disability with respect to the aforementioned outpatient applications, since the subject being monitored must either manually relay the information to the caregiver (such as by telephone, mail, or visit), or perform the analysis or interpretation themselves. Additionally, it is often desirable to perform more sophisticated (e.g. algorithmic) comparative or trend analysis of the subject's data, either with respect to prior data for that same subject, or data for other subjects. The lack of effective transmission modes in the prior art to some degree frustrates such analysis, since even if the subject has the facility to perform the analysis (e.g., PC or personal electronic device with the appropriate software), they will not necessarily be in possession of their own prior data, which may have accumulated via monitoring at a remote health care facility, or that for other similarly situated subjects.
Defibrillation
External defibrillators and defibrillation techniques are well known in the prior art. During external defibrillation, a strong current of short duration is administered across the thorax of the subject via paddles or electrodes to convert rapid non-pulsatile twitching of the heart ventricles, or ventricular fibrillation (VF), to a slower pulsatile rhythm that allows the heart to pump blood. Successful defibrillation stimulates cells by passing a given current intensity through the cells for a period of time. Typically, in the clinical environment, highly trained personnel determine if a defibrillation current (shock) should be administered. Outside of the clinical environment, personnel with little or even no training are often tasked with using portable or similar defibrillators, such as the well known automatic external defibrillator (AEDs), to analyze cardiac rhythms and, if appropriate, advise/deliver an electric shock.
A critical feature of defibrillators used by non-skilled personnel, and especially AEDs, is the accurate analysis of the electrocardiogram rhythm to determine if a defibrillation shock is necessary. As recommended by the American Heart Association (AHA) Task Force on AED, coarse VF (peak-to-peak amplitude>200 μV) and rapid ventricular tachycardia (VT) rhythms should be shocked, as they are almost always associated with a pulseless, unresponsive patient. Ventricular tachycardia refers to a rapid heartbeat (e.g., >100 beats per minute) originating in the right or left ventricle which prevents the ventricles from filling adequately with blood. This inadequate filling precludes the heart from pumping normally. VF should generally be detected and shocked with >90% sensitivity; rapid VT should generally be detected and shocked with >75% sensitivity.
Benign rhythms should also be identified, and not shocked with >95% specificity. Such benign rhythms include for example normal sinus rhythm, atrial fibrillation (rapid twitching of the atrium), sinus bradycardia (slow heartbeat), supraventricular ventricular tachycardia (VT originating outside of the ventricles, or SVT), heart block (delayed normal flow of electrical impulses), premature ventricular contractions, and idioventricular rhythms (slow, regular rhythms without coordination between the atria and ventricles). To maximize safety in the event the electrodes were misapplied, asystole (no electrical activity) should be detected and not shocked with >95% specificity. See, e.g., Kerber, et al., “Automatic external defibrillators for public access defibrillation: recommendations for specifying and reporting arrhythmia analysis algorithm performance, incorporating new waveforms, and enhancing safety”, Circ, 95:1677–1682, 1997.
Historically, arrhythmia detection has been performed using empirical methodology. For example, the Heartstream Forerunner manufactured by Philips Corporation, an exemplary prior art AED, uses a set of four indicators relating to ECG waveform shape and timing to provide the advisory function. These indicators are (i) rate, (ii) narrowness of the QRS complex, (iii) repeatability of QRS complexes, and (iv) amplitude. While such indicators are often sufficient to detect VF, ventricular and supraventricular tachycardia may not be always be easily distinguished. During typical AED usage outside the clinical environment, the patient is already unconscious, and SVT is very rarely presented. The Powerheart CRM, an automatic defibrillator manufactured by Cardiac Science, Inc., is used in the clinical setting. In an automatic defibrillator, the device automatically delivers a shock, rather than just providing an advisory to shock. For an automatic defibrillator, it is critical that SVT and pulsatile VT, during hemodynamic stability, be detected accurately and the patient not be shocked, as the patient may be conscious. Further, in the clinical environment, pulseless versus unstable ventricular tachycardias need to be distinguished, as pulseless VT is best treated with countershock, but unstable VT is best treated with synchronized cardioversion (a shock with lower energy level that is timed with the QRS complex). Accurate SVT and VT detection is difficult using empirical methodology since SVT and VT waveform shapes are highly variable. See, e.g., U.S. Pat. No. 6,480,734 entitled “Cardiac arrhythmia detector using ECG waveform-factor and its irregularity” issued Nov. 12, 2002, and U.S. Pat. No. 6,490,478 entitled “System and method for complexity analysis-based cardiac tachyarrhythmia detection” issued Dec. 3, 2002, both to Zhang, et al. and assigned to Cardiac Science.
Currently implemented AED and automatic defibrillator arrhythmia detection algorithms (such as those of the Philips and Cardiac Science approaches) are based on waveform time domain morphologies, with decision-making strategies based on absolute threshold criteria. These empirical methodologies decrease the sensitivity for detecting rhythms such as rapid ventricular tacchycardia.
Another prior art approach attempts to capitalize on the root reason for performing defibrillator arrhythmia analysis; i.e., to determine if insufficient cardiac output is present to warrant countershock therapy. See, e.g., Johnson, et al., “The transthoracic impedance cardiogram is a potential haemodynamic sensor for an automated external defibrillator,” Eur Heart J, 19:1879–1888, 1998 (hereinafter “Johnson”). The Johnson approach uses the impedance cardiogram as a hemodynamic sensor for AEDs. However, their approach is flawed in that, inter alia, it fails to consider or account for the effects of electrode configuration on the electric field distributions and hence the resolution of significant ICG waveform features. Specifically, the Johnson apparatus utilizes a single electrode configuration, which significantly decreases B and X point resolution. B and X point resolution were further degraded by ensemble averaging of dZ/dt, with respect to the R point.
Further, Johnson employs manually selected beats and dZ/dt features, which is at best cumbersome. Also, the dZ/dt features chosen in Johnson to classify shockable rhythms are based on the C point amplitude, dZ/dtMAX. This method is less than optimal, since C point amplitude has been demonstrated to be highly variable on a beat basis in both normal and cardiac subjects, due to respiration-synchronous fluctuations and other physiological factors. For example, in congestive heart failure (CHF) patients, standard deviations of greater than 100% have been demonstrated by the Assignee hereof. Thus, C point-related features are prone to significantly more physiologic variability than those associated with the B and X points. With such C point-related features, only 80% of the nonshockable rhythms (20/25) could be classified, well below the AHA recommendation of 95% specificity. Hence, under the prior art approach of Johnson (and others), a significant number of patients not requiring shocking are erroneously shocked or not shocked due to misclassification and related inaccuracies. Erroneous shocks can lead to severe burns, while failure to shock during fibrillation leads to death.
Based on the foregoing, there is a need for an improved apparatus and method for determining the need for, and applying, defibrillating current to a subject. Such improved apparatus and method ideally would provide the user with an accurate advisory of the need for shock, and be capable of automatic application of the shock when needed without further user intervention if so selected. This capability would permit users of literally any skill and training level to accurately and properly treat subjects displaying a variety of different cardiac conditions, such treatment including discrimination of those situations where shock is or is not required.